1. Field of the Invention
This invention relates to chucked and motor driven intramedullary bone machining tools, specifically to such tools used in the field of orthopedic surgery.
2. Description of the Prior Art
Orthopedic Surgeons commonly use methyl-methacrylate bone cement to secure prostheses to bone. To do so, the bone is commonly machined in such a way as to allow for stable cementation of the prosthesis to bone. In hip replacement surgery, this machining process includes the use of a series of reamers to cylindrically machine the endosteal canal of the femur. This is followed by the use of a broach to further shape the endosteal canal of the femur to match the shape of the implant.
Prosthetic hip joint replacement is often performed with the use of methyl-methacrylate bone cement. Bone cement is used to secure the artificial joint component to bone. A common method used for implantation of the femoral component of a hip replacement in humans is to cement the component to the bone. To do so requires the endosteal canal of the femur to be machined to a similar shape as that of the prosthesis. A series of bone reamers and broaches is used to create a cavity in the bone of similar shape to the implant, only larger. The larger size of the bone cavity relative to the prosthesis is to allow space between the two to place bone cement. The actual size difference between the machined cavity in the bone and the prosthetic implant will determine how thick the cement will be between the prosthesis and the bone.
The strength of the cement-to-bone bond has several variables which affect that bond. These variables affect this bond's ability to withstand repetitive loading and other stresses over a long term. These variables include the purity and porosity of the cement. Also included are any contaminants present at the cement-bone interface that might affect bonding. Another important variable is the bone surface topography to which the cement will be made to come in contact with and adhere to. The more porous and the more irregular the bone surface is, the greater the bond strength will be between the bone and the cement.
In the situation of primary hip replacement, no previous joint replacement has been performed on that hip. This is the situation in which bone porosity and bone surface topography are the most favorable to allow for a strong cement-to-bone bonding. In the situation of primary joint replacement, the surfaces of the bone are that of live vascularised bone tissue. These bony tissues have a great deal of porosity. The porosity arises from Haversian canals, which are neurovascular micro-canals communicating with the bone's surface. In addition, there is generally cancellous bone attached to the cortical bone in the endosteal canal in a bone which has not been previously operated upon. This cancellous bone provides interstices into which cement can fill and interlock. This interlocking of cement into cancellous bone increases the bond strength of the cement to the host bone.
In the situation of revision hip replacement, the bone surface is different than it is for primary hip replacement. In the revision situation, the bone surface is not favorable for allowing a strong cement-to-bone bonding. This is because bone undergoes a biological and mechanical response to a previously cemented implant. As part of this response, the cortical endosteal bone surfaces become sclerotic. Also, all cancellous bone in the endosteal canal in the region of the previous implant has been lost. In revision hip replacement surgery, a failed implant is removed and a new implant is replaced. The endosteal canal of a removed failed femoral component is smooth and sclerotic. The bone surfaces in this situation have very low porosity and very low surface roughness. For these reasons, the ability of bone cement to bond securely to this type of endosteal surface is greatly compromised. Much orthopedic research has been devoted to observations of patients with total hip replacements. This research reveals that in revision cemented hip replacement surgery in humans, the time from implantation of the cemented femoral component until the cement-to-bone bond fails and loosening occurs is much shorter than for analogous primary surgery of the same type. The reasons for significantly poorer results in the revision situation are as follows: In revision surgery the poor mechanical and physical qualities of the bone limit the cement-to-bone bond strength. The low porosity of the bone, which is smooth and sclerotic in a revision situation, limits the ability of cement to bond to the bone. Also the loss of cancellous bone in the revision state decreases the cement's ability to bond to the bone. For these reasons the cement-to-bone bond strength is considerably weaker in revision than in primary femoral component replacement, in human total hip replacement surgery. The long term follow-up of thousands of human hip replacement patients confirms this.
Throughout the useful life of a joint replacement implant, that implant will be subjected to many stresses. The stresses will be frequent and repetitive. Factors known to affect the ability of a cemented hip replacement to withstand these stresses are as follows:
(a) The strength of the cement-to-bone bond. The weaker this bond is, the shorter the time period will be from implantation until failure of fixation. Failure of fixation of cement to the bone will result in mechanical and clinical failure of the implant. PA1 (b) The strength of the cement-to-implant bond. This bond is generally stronger than the cement-to-bone bond. The cement-to-implant bond is not affected by the surgery being either primary or revision. PA1 (c) Mechanical factors of the implant. These include modulus of elasticity. Also included is the ability of the implant to withstand cyclical loading. This is measured as the number of load cycles to mechanical failure of the implant. PA1 (d) The thickness and uniformity of the cement mantle between the bone and the implant. This cement mantle secures the implant to the bone. The thicker and the more uniformly thick the cement mantle is, the better it will be able to withstand the stresses that are placed upon it. The better it is able to withstand those stresses, the longer the implant will remain in service. PA1 (e) The mechanical factors of the bone. These include surface roughness and porosity. These factors affect the strength and longevity of the cement-to-bone bond. PA1 (a) A high mismatch of modulus of elasticity between the implant and the bone will cause the patient to experience pain. The internal and external dimensions of the bone are a given. Since cement is not used in biological fixation hip replacement, a larger size femoral component must be used than had a cemented component been used. This is so because a cemented component would need to be smaller than the internal canal size of the femur, to allow room for the cement. Femoral components used are generally made from either of two materials; cobalt chromium alloy or titanium. Both of these materials have stiffness and moduli of elasticity considerably greater than that of bone. The thicker the implant is, the stiffer it will be. Thus, for a given size femur, there will be a higher mismatch of modulus of elasticity between bone and implant with a porous coated biological fixation implant than with a smaller cemented implant. This modulus of elasticity mismatch can produce symptoms of thigh pain when walking. There is a much higher incidence of thigh pain while walking in patients receiving porous coated biological fixation femoral implants then there is in patients receiving cemented femoral implants. This has been shown to be so in both human primary and revision hip replacement surgery, in multiple large studies done to date. PA1 (b) Bone ingrowth into the porous coating does not always occur. For a porous coated implant to become attached to bone in a stable way, host bone must grow into the pores. The frequency of bone ingrowth in-vivo is not 100%. It is on the order of 98% in primary femoral component human hip replacement surgery. The frequency of bone ingrowth in revision femoral component surgery is on the order of 92%. Failure of bone ingrowth can lead to a painful mechanical failure of the implant. PA1 (c) Weight bearing must be limited on the operative extremity. To minimize stress at the implant-bone interface and to maximize ingrowth potential, weight bearing is limited on the operative extremity for some time after surgery. This is temporary, but requires the patient to use ambulatory aids, such as crutches or a walker, for up to three months after surgery. With a cemented femoral stem, the patient could fully bear weight on that femur immediately. PA1 (d) A biological fixation type implant can migrate before ingrowth occurs. The bond of the porous coated implant to the bone depends upon the ingrowth of bone into the porous coating for strength and stability. Thus, the implant will not achieve optimal implant stability for six to twelve weeks after implantation. During this time period, if stresses on the implant exceed its preliminary press fit mechanical strength at its interface, the implant will move relative to the host bone. In the femur, the implant may change position vertically or rotationally. The implant may subside into the femoral canal. Subsidence will result in shortening of the operative limb and a limb length discrepancy may result. Such a shortening of the limb length at the hip will affect the soft tissue balance at the hip joint. The soft tissue balance provides the necessary mechanical forces to prevent hip dislocation. Thus a subsidence of the femoral stem may result in hip joint instability. This may predispose the patient to hip joint dislocations after surgery. In addition to vertical migration, it is possible for the prosthesis to make a rotational migration relative to the host bone. A rotatory migration of the femoral component will also affect stability of the hip. In addition, rotatory migration of the femoral component will affect the range of motion of the joint. In contrast, a cemented femoral component is stable immediately, and can not migrate relative to the femur. PA1 (e) The cost of manufacture is considerably higher to produce a porous coated component compared to a standard component. This is because an expensive extra step is required to apply the porous coating to the prosthesis. Femoral components that are implanted with cement are generally not porous coated.
The long term failure rate for a femoral prosthesis recemented at revision human hip replacement surgery is unacceptably high. For this reason, another method of fixing a prosthesis to bone has been developed. This other method is called biological fixation.
The most common method of fixing a femoral implant to bone without the use of cement is that of biological fixation. In this method, the femoral intramedullary canal is machined internally to the exact dimensions of the prosthesis to be implanted. A variety of implant sizes is available to facilitate this process. The prosthesis is coated at manufacture with a porous coating. The pore size is usually 200-400 microns in average size. Bone in contact with this type of implant surface will usually demonstrate some degree of bony ingrowth into the pores of the implant. This process usually takes six to twelve weeks from the time of implantation. When ingrowth occurs, it results in a strong mechanical interface between the bone and the implant. Using standard techniques, the long term success of human revision femoral component hip replacement surgery with biological fixation is significantly better than using cement for reimplantation. There are, however, a number of disadvantages of using a porous coated biological fixation type implant as opposed to a standard cemented one in human hip replacement surgery as follows:
To the inventor's knowledge, no one has previously disclosed a method or apparatus for placing grooves in a bone for the purpose of improving cement-to-bone bonding. U.S. Pat. No. 4,550,788 to Park (1985) discloses a roof bolt hole groover. This is a drilling device for use in bolting together a discontinuous rock mass in order to provide a safe mine roof. The grooving portion of this device includes at least two groove bits that reciprocate pivotally for the purpose of scoring the surface of the hole being drilled. This is primarily a drilling device, and secondarily a scoring device. This device has a grooving mechanism coupled to a drilling mechanism. The two mechanisms are coupled by a series of gears so that the two operations are done simultaneously. The purpose of this device is to score a hole while it is being drilled. This device would be impractical to use to create multiple grooves inside the diaphysis of a bone. This device is not easily disassembled for cleaning purposes, as would be required of a medical device suitable for grooving of bone. The reciprocating mechanism for the grooving bits is a significant disadvantage to its application as a groover. The bits reciprocate toward and away from the main axis of the device while it is in operation. This will result in a tendency for the device to move axially at the point of operation in which the groove bits are maximally retracted. This, in turn, would result in inaccuracies in grooving, and groove widening. Thus this device is not conducive to making finely controlled grooves in human bone during surgery, where accuracy is of great importance. This device is not intended to create circumferential grooves while remaining stationary in an axial direction. It will not work well to cut multiple grooves of constant width, depth, or spacing. The non-reproducibility of exact grooving patterns would significantly limit its usefulness in the medical field for cutting grooves in live human bone. This device contains an end cutting drill. This drill could damage the bone or material being machined in unintended ways if attempting to use the grooving mechanism to cut thin uninterupted grooves. This device has no intended medical uses.
U.S. Pat. No. 3,472,229 to Kuntscher (1969) discloses an instrument for cutting or severing a bone from the inside thereof. This device consists of a surgical saw having an elongated flexible shank and drive support. This device is intended for severing a bone from the inside. This device could be used to cut grooves in a bone. However, to do so would be quite difficult. There would be minimal if any visibility at the cutting site. The operation of this device would most likely require x-ray flouroscopic monitoring. The resultant radiation exposure would have risks to the patient and nearby personnel. The cutting saw is such that it is generally at an acute angle to the bone. This acute angle would make circumferential grooving of the bone's endosteal canal very time consuming and difficult. The saw will have a great tendency to bounce off of the bony surface while cutting. This is because of the long lever arm controlling the saw. This device will work poorly to cut multiple circumferential grooves of constant width, depth, or spacing. The non-reproducibility of exact grooving patterns is a significant disadvantage of this device when used for grooving a bone canal.
U.S. Pat. No. 3,435,729 to Toth (1967) discloses a hole grooving device. The purpose of this device is for toughening the inner formed finger holes of a bowling ball. This device contains a pair of semi-cylindrical body members containing the grooving mechanism, and an elongated rod extending radially from the body members which is the mechanism of deployment of the cutting member of this device. This device has several disadvantages which make this a poor if not useless tool for grooving the intramedullary canal of a long bone. One major disadvantage of this device is that the deployment mechanism results in the two semi-cylindrical body members moving apart from each other and pushing against the walls of the hole being machined. This creates excess friction between the device and the hole being machined. This will also tend to deform the hole being machined. In a bone, this could easily result in a fracture of the bone if too much pressure is applied against the inner walls of the bone canal by the body of this device. Another disadvantage of this device is that the elongated rod extending radially from the body of this device will impinge upon bone and soft tissues if used to groove the canal of a long bone during live surgery. In the human femur during hip replacement surgery, this rod would impinge upon the greater trochanter, making the device useless for producing circumferential grooves in the intramedullary canal of the proximal femur.
U.S. Pat. No. 5,387,218 to Meswania (1995) discloses a surgical instrument for shaping bone. This device functions as a motorized broach to machine the intramedullary canal of a bone. This device however, does not have a retractable cutting member. It therefore can not produce grooves in the diaphyseal portion of an intramedullary bone canal. This device is designed to machine and shape the bone, and not to machine circumferential grooves into the bone's surface. Since the cutting surface is only on one side of this device, and since the entire device can only move axially in the bone, this device can not be used to produce circumferential grooves in any portion of the bone canal.